Bremsstrahlung target for radiation therapy system

ABSTRACT

Described herein is a medical linear accelerator including an accelerator target structure constructed of a material having a thickness of less than 0.2 radiation lengths, and an accelerator structure to receive an electromagnetic wave and generate an output therapy dose rate of electrons having a beam energy between 4-25 mega-electronvolts (MeV).

RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.15/288,867, filed Oct. 7, 2016, which claims the benefit of U.S.Provisional Application No. 62/239,608, filed Oct. 9, 2015, the entirecontent of which are incorporated herein by reference.

TECHNICAL FIELD

Embodiments of the invention relate to a Bremsstrahlung target forradiation therapy.

BACKGROUND

Accelerator-based radiation therapy typically generates a high energyX-ray beam via bremsstrahlung (“braking radiation”). A relativisticelectron beam is incident on a target material of high atomic number(“high Z”). The electrons are deflected (accelerated) by electromagneticinteractions with the target nuclei, causing emission of high energyphotons. Some of these photons have enough energy to createelectron-positron pairs, which then interact with target nuclei to emitmore photons. The result is an “electromagnetic shower” or“electromagnetic cascade” of electrons, positrons, and photons. Anyelectrons which escape the target are typically eliminated from thetherapy beam by an electron absorber made of low Z material (e.g.aluminum, carbon).

BRIEF DESCRIPTION OF THE DRAWINGS

Embodiments of the present invention are illustrated by way of example,and not by way of limitation, in the figures of the accompanyingdrawings.

FIG. 1 illustrates an image-guided radiation treatment system, inaccordance with embodiments of the present invention.

FIG. 2 illustrates a cross-section of a linear accelerator in accordancewith one embodiment of the present invention.

FIG. 3 illustrates an electron beam accelerated into a linearaccelerator target to produce x-rays in accordance with an embodiment ofthe present invention.

FIGS. 4A and 4B illustrate a linear accelerator target structure inaccordance with embodiments of the present invention.

FIG. 5 illustrates a gantry based image-guided radiation treatmentsystem, in accordance with embodiments of the present invention.

DETAILED DESCRIPTION

Embodiments of the present invention relate to a Bremsstrahlung targetin a medical linear accelerator (LINAC) for radiation therapy. In anexemplary LINAC, electrons injected into an accelerator structure of theLINAC by an electron gun are accelerated and directed along theaccelerator structure using the electric and magnetic field componentsof an electromagnetic wave that is coupled into the acceleratorstructure. The electromagnetic wave may be coupled into the acceleratorstructure from an amplifier, such as a klystron, or an oscillator, suchas a magnetron. As the electrons traverse the accelerator structure,they are directed and accelerated by forces exerted on the electrons bythe electric and magnetic field components of the electromagnetic waveto produce a high-energy electron beam. In some embodiments, thedirecting of the electrons may be assisted by static magnetic fieldsfrom solenoids, dipoles, quadrupoles or combined-function magnets. Inother embodiments, the directing of the elections may not be assisted bystatic magnetic fields. The electron beam from the accelerator structuremay be directed at an x-ray emitting target (referred to as aBremsstrahlung target) to generate x-rays. Although embodiments of thepresent invention may be described using a traveling wave LINAC, itshould be noted that embodiments of the present invention may also beutilized in any electron accelerator capable of reachingmega-electronvolt (MeV) beam energies. Examples of electron acceleratorscapable of reaching MeV beam energies include, but are not limited to,standing wave radio frequency (RF) LINACs, betatrons, dynamitrons,rhodotrons, synchrotrons and the like.

The x-ray emitting target is comprised of a material of high atomicnumber (“high Z”). As the thickness of the high Z target material isincreased, the amount of radiation produced is increased. However, thehigh Z target material also absorbs radiation. If the thickness of thetarget is too thick, the result will be a decrease in total radiationflux. Therefore, in accordance with embodiments of the presentinvention, a the LINAC is designed with a particular thickness for thehigh Z target material that balances photon production and absorption tomaximize the total photon flux from the target, as will be discussed inmore detail in FIG. 3 below.

The optimal thickness for the LINAC target may be expressed in terms ofradiation length of the incident electrons in the high Z material. Theradiation length is the mean distance over which a high-energy electronloses all but 1/e of its energy by bremsstrahlung. In one embodiment,the thickness of the target is in a range of 0.25 to 2 radiation lengthsfor bremsstrahlung targets without much variation of photon yield overthis range of thicknesses. For example, the radiation length for a highenergy electron in solid tungsten is approximately 3.5 millimeters (mm).Therefore, tungsten targets optimized for producing X-rays fromrelativistic electrons may be 0.9 to 7 mm in thickness.

As the electrons lose energy in the LINAC target, the energy density ofthe target increases. In order to prevent failure of the target it mustbe properly cooled. One method of cooling a LINAC target is to transferenergy from the target to a heat sink material with high thermalconductivity (e.g. copper) which is actively cooled with water. However,as therapy dose rates increase and as electron spot sizes are reduced todiminish penumbra, the energy density in the high Z material continuesto increase, making cooling the high Z material difficult even whenusing the two-layer structure described above. A result may be theoverheating and failure of a target.

An embodiment of the present invention resolves the cooling issuedescribed above by minimizing the thickness of the high Z target. Byminimizing the thickness of the LINAC target, the two-layer coolingstructure is able to sufficiently cool the target at the increasedtherapy dose rates and electron spot sizes. The result being a LINACcapable of administering higher dose rates without overheating andfailure of the high Z target.

The x-ray emitting target may be comprised of a material having athermal conductivity of about 50 watts per meter-Kelvin (W/m·K) orhigher in some implementations. For example, the x-ray emitting targetmay be comprised of a high Z target material having a thermalconductivity of about 50-400 W/m·K. Cooling of the x-ray emitting targetmay be facilitated by using a target material that has an increasedthermal conductivity (e.g., a thermal conductivity of about 50 W/m·K orabove). In other implementations, the x-ray emitting target may becomprised of a material having a thermal conductivity of about 8-400W/m·K. Table 1 below lists some possible materials for the x-rayemitting target.

TABLE 1 Thermal Conductivities for Candidate Target Materials MaterialThermal Conductivity (W/m · K) Aluminum 204 Copper 386 Tantalum 54Tungsten 165 W25Re 60 Rhenium 71 Platinum 73 Gold 315 Mercury 8 Lead 35Uranium 24

FIG. 1 illustrates an image-guided radiation treatment system, inaccordance with embodiments of the present invention. In the illustratedembodiment, the radiation treatment system 100 includes a LINAC 101 thatacts as a radiation treatment source. In one embodiment, the LINAC 101may be a standing-wave LINAC. In an alternative embodiment, the LINAC101 may be a traveling wave LINAC. In one embodiment, the LINAC 101 ismounted on the end of a robotic arm 102. In another embodiment, theLINAC 101 may be mounted on a gantry based system as illustrated in FIG.5. LINAC 101 delivers one or more radiation treatment beams to atreatment target 120 within patient 125. In one embodiment the LINAC 101may be an S-Band LINAC. In other embodiments the LINAC 101 may be aC-Band, X-Band or L-Band LINAC. FIG. 1 will be discussed in more detailbelow.

FIG. 2 illustrates a cross-section of a LINAC in accordance with oneembodiment of the present invention. In the illustrated embodiment, theLINAC 101 includes an electron gun 210. An example electron gun 210includes an anode, a grid, a cathode and a filament. The filament isheated to cause the cathode to release electrons, which are acceleratedaway from the cathode and towards the anode at a high speed. The anodecan focus the stream of emitted electrons into a beam of a controlleddiameter. The grid can be positioned between the anode and the cathode.

The electromagnetic wave source 220 is a linear-beam vacuum tube thatreceives the electron beam from the electron gun 210 and generates highpower electromagnetic waves (carrier waves). In one embodiment, theelectromagnetic wave source 220 may be a magnetron. In anotherembodiment, the electromagnetic wave source 220 may be a klystron. Theelectromagnetic wave source 220 provides the driving force that powersthe LINAC 101. The electron tube 220 coherently amplifies the inputsignal to output high power electromagnetic waves that have preciselycontrolled amplitude, frequency and input to output phase in the LINACaccelerator structure.

High power electromagnetic waves are injected into the acceleratorstructure 230 from the electron tube 220. The electrons enter theaccelerator structure 230 and are typically bunched in the first fewcells of the accelerator structure 230. The accelerator structure 230 isa vacuum tube that includes a sequence of tuned cavities separated byirises. The tuned cavities of the accelerator structure 230 are boundedby conducting materials such as copper to keep the energy of the highpower electromagnetic waves from radiating away from the acceleratorstructure 230.

The tuned cavities are configured to manage the distribution ofelectromagnetic fields within the accelerator structure 230 anddistribution of the electrons within the electron beam 260. The highpower electromagnetic waves travel at approximately the same speed asthe bunched electrons so that the electrons experience an acceleratingelectric field continuously. In the first portion of the LINAC 101, eachsuccessive cavity is longer than its predecessor to account for theincreasing particle speed. The basic design criterion is that the phasevelocity of the electromagnetic waves matches the particle velocity atthe locations of the accelerator structure 230 where accelerationoccurs.

Once the electron beam 260 has been accelerated by the acceleratorstructure 220 and passes through the final accelerator cavity 270, itcan be directed at target 240 (e.g., constructed of a material such as atungsten or copper) that is located at the end of accelerator structure220. In one embodiment, the target 240 may be coupled to a heat sink 250to aid in the cooling of the target 240. The target 240 and the heatsink 250 will be discussed in more detail in FIG. 4 below. Thebombardment of the target 240 by the electron beam 260 generates a beamof x-rays (as discussed in FIG. 3 below). The electrons can beaccelerated to different energies before striking the target 240. In oneembodiment, the electron beam 260 may have a beam energy in the range of4 to 25 MeV.

FIG. 3 illustrates an electron beam 260 accelerated into a target 240 toproduce x-rays. In the illustrated embodiment, an electron beam 260 isaccelerated into a target 240, causing the emission of x-rays 310.

LINAC 101 only allows a small fraction of the total photons to passthrough the system. Only those photons having small angles 320 withrespect to the incident electron beam are kept, while photons havinglarge angles 330 are absorbed by a collimation system 340. As theincident electrons interact with the target material, they scatter andfill an increasing solid angle. Most of the acceptable photons originatein the early interactions of an incident electron with the target 240.Following the early interactions, the incident electrons will typicallyhave an angle that is too large to produce photons that will be acceptedby the collimation system 340. For example, a 6 MeV electron beam willdevelop a route mean square (RMS) angular spread of approximately 15degrees after passing through 0.01 radiation length of target material.After passing through 0.1 radiation length of target material, the RMSangular spread increases to approximately 45 degrees. While theelectrons that are scattered throughout the large angles will continueto produce photons, very few of these photons will be accepted by thecollimation system 340 and will contribute to therapy.

When the electrons have passed through a sufficient target thicknessthat the electron scattering angle exceeds the natural bremsstrahlungangle, the photon angular distribution will be dominated by the electronscattering angle. The electron scattering angle grows corresponding tothe square root of the target thickness, resulting in the solid angleover which photons are emitted growing as the target thicknessincreases. The bremsstrahlung photon production also grows as the targetthickness increases. Therefore, the photon density into a smallcollimated angle is roughly constant after the electron scattering angleexceeds the natural bremsstrahlung angle. In the previously describedexample, for a 6 MeV electron beam the electron scattering angle exceedsthe natural bremsstrahlung angle at approximately 0.01 radiation length.As such, target 240 may be made thinner than conventional bremsstrahlungtargets without loss of photon flux. Furthermore, there may be aresultant increase in photon flux due to lower photon absorption.Another advantage of the present embodiment is that a thinner targetproduces fewer total photons. Therefore, there will be less scatteredradiation into electronics and other components, enabling radiationshielding thickness to be decreased.

FIG. 4A illustrates one embodiment of a LINAC target structure. Theillustrated embodiment of the LINAC target structure 400 includesbonding material 410, an x-ray emitting target 240 having a targetthickness 430, bonding material 420 and a heat sink 250. In oneembodiment, bonding materials 410 and 420 are brazing alloys used tobraze the x-ray emitting target 240 to the heat sink 250. Examples ofbrazing alloys include, but are not limited to, aluminum, copper, brass,bronze, nickel, silver and the like. However, one skilled in the artwould recognize the present invention may utilize other forms of bondingincluding, but not limited to, soldering, explosion bonding, diffusionbonding and the like. In an alternative embodiment, the LINAC targetstructure 400 does not include brazing material 410.

The x-ray emitting target 240 may be constructed from a metal materialsuch as tungsten or a tungsten alloy. However, in alternativeembodiments the LINAC target 240 may be comprised of any material havingan atomic number greater than or equal to 40 (i.e. a high Z material).Examples of alternative materials for target 240 include, but are notlimited to, tantalum, rhenium, platinum, gold, liquid mercury, liquidlead, uranium or any alloys or mixtures of high Z materials. In someembodiments the LINAC target 240 may be a circular disk. In otherembodiments the LINAC target 240 may be a foil. The LINAC target 240 hasa thickness 430 range of 0.01 to 0.2 radiation lengths. For example, inthe present embodiment using a tungsten x-ray emitting target 240 havinga radiation length of 3.5 mm, the target 240 would have an actualthickness 430 range of 0.035-0.7 mm. It should be noted that, while thethickness 430 range of 0.01 to 0.2 radiation lengths remains constant,the actual thickness 430 range of the target 240 will vary based on thetarget material used. In some embodiments the diameter of the target 240may be the same diameter as the electron beam from the acceleratorstructure. In other embodiments the diameter of the target 240 may belarger than the electron beam diameter.

In the embodiment illustrated in FIG. 4A, the target 240 is located in arecessed portion of heat sink 250. FIG. 4B is a side profile view of analternative embodiment where the heat sink 250 may not include arecessed portion, in which case the target 240 may be coupled to thesurface of heat sink 250 by bonding material 410. In other embodiments,the LINAC target structure 400 may not include a heat sink 250 orbrazing alloys 410 and 420 and be comprised solely of the target 240.The heat sink 250 is constructed from a material having a high thermalconductivity to aid in the cooling of the target 240. Examples ofmaterials having a high thermal conductivity include, but are notlimited to, copper, aluminum and brass.

Referring back to FIG. 1 illustrating configurations of image-guidedradiation treatment system 100. In one embodiment, the LINAC 101 aremounted on the end of a robotic arm 102 having multiple (e.g., 5 ormore) degrees of freedom in order to position the LINAC 101 to irradiatea pathological anatomy (e.g., treatment target 120 within patient 125)with beams delivered from many angles, in many planes, in an operatingvolume around a patient 125. Treatment may involve beam paths with asingle isocenter, multiple isocenters, or with a non-isocentricapproach. Alternatively, other types of image guided radiation treatment(IGRT) systems may be used. In one alternative embodiment, the LINAC 101may be mounted on a gantry based system, for example, as illustrated inFIG. 5

The LINAC 101 may be positioned at multiple different nodes (predefinedpositions at which the LINAC 101 stops and radiation may be delivered)during treatment by moving the robotic arm 135. At the nodes, the LINAC101 can deliver one or more radiation treatment beams to a treatmenttarget 120. The nodes may be arranged in an approximately sphericaldistribution about a patient. The particular number of nodes and thenumber of treatment beams applied at each node may vary as a function ofthe location and type of pathological anatomy to be treated.

The radiation treatment system 100, in accordance with one embodiment ofthe present invention, includes an imaging system 165 having a processor130 connected with x-ray sources 103A and 103B and fixed x-ray detectors104A and 104B. Alternatively, the x-ray sources 103A, 103B and/or x-raydetectors 104A, 104B may be mobile, in which case they may berepositioned to maintain alignment with the treatment target 120 withinpatient 125, or alternatively to image the treatment target 120 fromdifferent orientations or to acquire many x-ray images and reconstruct athree-dimensional (3D) cone-beam CT. In one embodiment the x-ray sourcesare not point sources, but rather x-ray source arrays, as would beappreciated by the skilled artisan. In one embodiment, LINAC 101 servesas an imaging source (whether gantry or robot mounted), where the LINAC101 power level is reduced to acceptable levels for imaging.

Imaging system 165 may perform computed tomography (CT) such as conebeam CT, and images generated by imaging system 165 may betwo-dimensional (2D) or three-dimensional (3D). The two x-ray sources103A and 103B may be mounted in fixed positions on the ceiling of anoperating room and may be aligned to project x-ray imaging beams fromtwo different angular positions (e.g., separated by 90 degrees) tointersect at a machine isocenter (referred to herein as a treatmentcenter, which provides a reference point for positioning the patient ona treatment couch 106 during treatment) and to illuminate imaging planesof respective detectors 104A and 104B after passing through the patient125. In one embodiment, imaging system 165 provides stereoscopic imagingof the treatment target 120 within patient 125 and the surroundingvolume of interest (VOI). In other embodiments, imaging system 165 mayinclude more or less than two x-ray sources and more or less than twodetectors, and any of the detectors may be movable rather than fixed. Inyet other embodiments, the positions of the x-ray sources and thedetectors may be interchanged. Detectors 104A and 104B may be fabricatedfrom a scintillating material that converts the x-rays to visible light(e.g., amorphous silicon), and an array of CMOS (complementary metaloxide silicon) or CCD (charge-coupled device) imaging cells that convertthe light to a digital image that can be compared with a reference imageduring an image registration process that transforms a coordinate systemof the digital image to a coordinate system of the reference image, asis well known to the skilled artisan. The reference image may be, forexample, a digitally reconstructed radiograph (DRR), which is a virtualx-ray image that is generated from a 3D CT image based on simulating thex-ray image formation process by casting rays through the CT image.

FIG. 5 illustrates one embodiment of a gantry based (isocentric)intensity modulated radiotherapy (IMRT) system 500. In a gantry basedsystem 500, a radiation source (e.g., a LINAC 101) having a headassembly 501 are mounted on the gantry in such a way that they rotate ina plane corresponding to an axial slice of the patient. Radiation isthen delivered from several positions on the circular plane of rotation.In IMRT, the shape of the radiation beam is defined by a multi-leafcollimator (MLC) that allows portions of the beam to be blocked, so thatthe remaining beam incident on the patient has a pre-defined shape. Theresulting system generates arbitrarily shaped radiation beams thatintersect each other at the isocenter to deliver a dose distribution tothe treatment target 120. In one embodiment, the gantry based system 500may be a C-arm based system.

Unless stated otherwise as apparent from the foregoing discussion, itwill be appreciated that terms such as “processing,” “computing,”“generating,” “comparing” “determining,” “calculating,” “performing,”“identifying,” or the like may refer to the actions and processes of acomputer system, or similar electronic computing device, thatmanipulates and transforms data represented as physical (e.g.,electronic) quantities within the computer system's registers andmemories into other data similarly represented as physical within thecomputer system memories or registers or other such information storageor display devices. Embodiments of the methods described herein may beimplemented using computer software. If written in a programminglanguage conforming to a recognized standard, sequences of instructionsdesigned to implement the methods can be compiled for execution on avariety of hardware platforms and for interface to a variety ofoperating systems. In addition, embodiments of the present invention arenot described with reference to any particular programming language. Itwill be appreciated that a variety of programming languages may be usedto implement embodiments of the present invention.

It should be noted that the methods and apparatus described herein arenot limited to use only with medical diagnostic imaging and treatment.In alternative embodiments, the methods and apparatus herein may be usedin applications outside of the medical technology field, such asindustrial imaging and non-destructive testing of materials. In suchapplications, for example, “treatment” may refer generally to theeffectuation of an operation controlled by the treatment planningsystem, such as the application of a beam (e.g., radiation, acoustic,etc.) and “target” may refer to a non-anatomical object or area.

The above description of illustrated embodiments of the invention,including what is described in the Abstract, is not intended to beexhaustive or to limit the invention to the precise forms disclosed.While specific embodiments of, and examples for, the invention aredescribed herein for illustrative purposes, various equivalentmodifications are possible within the scope of the invention, as thoseskilled in the relevant art will recognize. The words “example” or“exemplary” are used herein to mean serving as an example, instance, orillustration. Any aspect or design described herein as “example” or“exemplary” is not necessarily to be construed as preferred oradvantageous over other aspects or designs. Rather, use of the words“example” or “exemplary” is intended to present concepts in a concretefashion. As used in this application, the term “or” is intended to meanan inclusive “or” rather than an exclusive “or”. That is, unlessspecified otherwise, or clear from context, “X includes A or B” isintended to mean any of the natural inclusive permutations. That is, ifX includes A; X includes B; or X includes both A and B, then “X includesA or B” is satisfied under any of the foregoing instances. In addition,the articles “a” and “an” as used in this application and the appendedclaims should generally be construed to mean “one or more” unlessspecified otherwise or clear from context to be directed to a singularform. Moreover, use of the term “an embodiment” or “one embodiment” or“an implementation” or “one implementation” throughout is not intendedto mean the same embodiment or implementation unless described as such.Furthermore, the terms “first,” “second,” “third,” “fourth,” etc. asused herein are meant as labels to distinguish among different elementsand may not necessarily have an ordinal meaning according to theirnumerical designation.

In the foregoing specification, the invention has been described withreference to specific exemplary embodiments thereof. It will, however,be evident that various modifications and changes may be made theretowithout departing from the broader spirit and scope of the invention asset forth in the appended claims. The specification and drawings are,accordingly, to be regarded in an illustrative sense rather than arestrictive sense.

What is claimed is:
 1. A medical linear accelerator comprising: anelectromagnetic wave source to receive a beam current of electrons andto generate an electromagnetic wave; an accelerator structure to receivethe electromagnetic wave and to generate an output therapy dose rate ofelectrons having a beam energy between 4-25 mega-electronvolts (MeV),wherein the accelerator is a standing wave accelerator; and anaccelerator target structure to receive the output dose of electronscomprised of: an x-ray emitting target to emit x-rays in response toreceiving the output dose rate of electrons, wherein the x-ray emittingtarget is less than 0.2 radiation lengths.
 2. The medical linearaccelerator of claim 1, wherein the thickness of the x-ray emittingtarget is between 0.01 and 0.2 radiation lengths.
 3. The medical linearaccelerator of claim 1, wherein the accelerator is coupled to a roboticarm.
 4. The medical linear accelerator of claim 1, wherein theaccelerator is coupled to a gantry.
 5. The medical linear accelerator ofclaim 1, wherein the standing wave accelerator is an X-Band, C-Band,S-Band or L-Band linear accelerator.
 6. The medical linear acceleratorof claim 1, wherein the electromagnetic wave source is a magnetron. 7.The medical linear accelerator of claim 1, wherein the acceleratortarget structure is further comprised of a heat sink coupled to thex-ray emitting target.